The invention relates to an actively shielded, cylindrical gradient coil system for use in an MR (=magnetic resonance) spectrometer with a main field magnet, which generates a main magnetic field aligned in the direction of a z-axis, wherein, when current flows in one of the measurement volumes through which the z-axis passes, the gradient coil system generates a Z-gradient field, whose passage through zero is located at the center of the measurement volume, and wherein the gradient coil system has at least one main gradient coil and at least one active shielding coil, wherein the main gradient coil is constructed from at least two cylindrical partial coil systems axially spaced from one another in the z-direction by a length L1 and symmetrically with respect to the center of the measurement volume, the axes of said cylindrical partial coil systems extending collinearly with the z-axis, wherein the cylindrical partial coil systems are at least partially constructed from electrical conductor sections wound with a maximum outer radius R1gradientoutmax around the z-axis, wherein at least one of the active shielding coils is constructed from electrical conductors on a minimum inner radius R1shieldinmin around the z-axis, and wherein R/shie/dinmin>R1gradientoutmax. Such a gradient coil system for an imaging NMR apparatus is disclosed for example by U.S. Pat. No. 5,296,810.
A modern nuclear magnetic resonance (NMR) spectrometer consists of an electromagnet for generating a strong, static magnetic field, a shim system for homogenizing the static magnetic field and an NMR probe, comprising at least one transmit and/or receive coil system for transmitting RF pulses and receiving the signals, a measurement sample, and a gradient coil system for generating pulsed field gradients. Moreover, the NMR spectrometer comprises the necessary apparatus for generating and detecting electrical signals, which are generated and/or detected in the aforementioned components.
Most modern NMR probes contain actively shielded gradient coil systems for generating a Z-gradient field, in rare cases additionally for generating X-, Y-, and Z-gradient fields. The active shielding is necessary because many NMR pulse sequences require fast switching of the gradient fields, which, in the case of unshielded gradient coil systems, results in induction of eddy currents in the surrounding metallic structures (in particular, in the outer casing of the probes, the former for the shim system, and the various metallic elements of the superconductive magnet systems, some of which are cryogenically cooled, and the cryogenic shim systems). The objective of active shielding is to reduce eddy currents to a minimum and to reduce the measurement artifacts that are caused by the remaining eddy currents in the measurement volume. These measurement artifacts are grouped under the term “recovery characteristics” of the gradient coil system and comprise both phase and amplitude errors of the NMR lines to be received.
In the prior art, actively shielded gradient coil systems are usually manufactured in such a way that the turns of the gradient coils and the associated shielding coils are located at two different radii for each gradient. The turns on the inner radius are used for generating the gradient field while the outer radius is responsible for shielding the gradient fields toward the outside. In some cases, certain linearization tasks can be performed by the turns on the outer diameter and part of the shielding is located in the outermost axial regions of the inner cylinder.
This design principle offers great advantages in the manufacturing and calculation of gradient coils. In particular, manufacturing is simplified because when constructing gradient coils from cut electrically conductive tubes, foils, metal sheets, PCB material, or coating and structuring on cylindrical substrates, only two tube-shaped objects have to be aligned with each other in each case. In the case of multiple-layered gradient coils, additional radial and axial positioning has to be performed, which usually reduces the yield during manufacturing.
In particular, in Z-gradient systems, it is possible to choose a design for the gradient in which the distribution of the electrical conductors of the main gradient coil exhibits axial spacing, symmetrically with respect to the center of the measurement volume. Depending on the conceptual design, this can also be applied to active gradient shielding. Examples of the allocation of such gradients are provided in, for example, U.S. Pat. No. 4,733,189.
The turns of the gradients are usually connected in series, to be able to ensure a constant current through all conductors. Parallel connections would result in fluctuations of the gradient field by the variation of the currents of different partial coils, in particular, when the temperatures of the individual conductors vary during operation due to inhomogeneous cooling, differing conductor lengths, and the resulting resistances, etc. However, parallel connections are also used in practice, which demand greater effort in the generation and regulation of the gradient currents. The usual way of making an electrical connection between the two symmetrical halves of the main gradient coil and the active shielding coil is to route them through the central region in which they are spaced, wherein the conductor for connection can extend either coaxially with the cylinder axis or along any curve. Examples of curved profiles are provided in U.S. Pat. No. 7,109,712 B2 or in U.S. Pat. No. 6,456,076 B1.
Examples of straight profiles are given, for example, in U.S. Pat. No. 5,296,810 cited in the introduction or in U.S. Pat. No. 6,456,076 B1 or in U.S. Pat. No. 7,109,712 B2, documents cited as prior art.
Besides the tube-shaped gradient coil system, the prior art also comprises other geometries in which the turns of the main and/or shielding gradient coils are located on surfaces that are more complex:
U.S. Pat. No. 5,512,828 discloses gradients that consist of one main gradient coil and one active shielding coil, wherein the distance between the two coils is greater in the regions further from the center than in the regions near the center.
U.S. Pat. No. 5,939,882 discloses a gradient coil system in which the gradient coils do not occupy the entire space. However, this is a modification of a biplanar gradient coil system on curved surfaces and not a cylindrical gradient coil system. In the region of the RF coils, occupancy of the space by gradient coils is provided at least for part of the xy-plane.
U.S. Pat. No. 6,933,723 B2 discloses a gradient coil system in which the main gradient coil is relocated in the region of the RF coil. Herein, 218/221 represent the symmetry axes x and z, 211 the magnet for generating a static magnetic field, 212 the active shielding coil of the gradient coil system, 213/213′ the main gradient coil of the gradient coil system, and 219 the RF coil. Since neither the description nor the illustrations describe an RF shielding for limiting the volume accessible to the RF system, it must be assumed that there is no RF shielding in this configuration.
The same is disclosed in U.S. Pat. No. 7,852,083 B2; here, however, RF shielding is explicitly described, which gives the RF coil system a larger volume due to the set back part of the main gradient coil and thus increases the performance of the RF coil system (or permits a greater distance between the gradient coil and the shielding coil for the gradient in the region outside the central region and thus improves the efficiency of the gradient coils as compared with a conventional gradient coil system). The RF shielding is provided in the form of a main gradient coil, which, in the region of the RF coils, has a larger radius r2>r1 than outside the active RF region.
Common to both documents, U.S. Pat. No. 7,852,083 B2 and U.S. Pat. No. 6,933,723 B2, is the fact that the gradient coils each extend along the full length of the z-axis or even have an overlapping region in which turns of the gradient coils exist on two radii. There is no axial region, in which the main gradient coils have no turns.
U.S. Pat. No. 7,057,391 B1 discloses a magnet system with integrated gradients (3) and an RF coil (4), in which the gradient and the coil are set into a recess in the magnet. Here, the aim is to utilize the “unused” space for efficient generation of the static magnetic field. Gradient, RF coil, and possibly the RF shielding seem to have been manufactured cylindrically.
U.S. Pat. No. 6,154,110 discloses a gradient coil system for open MRI magnets in which the gradient coils and the shielding coils are discontinued in a central region. RF shielding cannot be mounted in that same region as otherwise the open system would be closed by the RF shielding. In principle, this is a modification of biplanar gradient coil systems.
U.S. Pat. No. 5,600,245 discloses local gradient coil systems which surround the RF coil and leave an opening free in the region of the RF coil. However, these gradient coils are not actively shielded coils and can only function in conjunction with a main gradient coil and are only used to strengthen the gradient field locally.
U.S. Pat. No. 5,406,204 discloses a gradient coil system that contains RF shielding, which in certain embodiments, is disposed on different radii. The turns of the Z-gradient coil in one embodiment are installed in grooves of a former. The RF shielding is either installed on the surface of the former and in the grooves, outside the Z gradient coil, or both in the grooves and outside the Z gradient coil.
The depth of the grooves more or less corresponds to the thickness of the Z-gradient turns. It is pointed out that the outside surface of the RF shielding, together with the outside surface of the Z gradient turns, should constitute a largely even surface for accommodating the X- and Y-gradient turns.
Compared with the large diameters of the gradient coil system (60-90 mm), the thickness of the gradient turns can be considered negligible. This can be seen, in particular, in FIGS. 1 to 3 of U.S. Pat. No. 5,406,204: The illustrated graduation of the RF shielding in the grooves of the Z-gradient turns extends only in the range of approx. 0.5 to 1% of its radius. Due to this slight variation of the radius, it is explicitly explained that the grooves are irrelevant to the functionality of the invention and that the Z-gradient could be laminated onto a continuous RF shielding of constant radius without any substantial loss of performance. It seems that the RF shielding follows the shape of the grooves for manufacturing reasons rather than for performance reasons.
The document does not clearly state how the electrical connections between the individual turns of the Z gradient coil are implemented. However, it is stated that the X-, Y-, and Z-gradient coils are preferably designed according to patent application Ser. No. 07/942,521. This is the application on which the published application U.S. Pat. No. 5,296,810 cited in the introduction is based, in which a series connection of the turns for a Z-gradient is also shown to pass through the central region. It can therefore be assumed that in U.S. Pat. No. 5,406,204, electrical conductors exist across the entire length of the gradient coil system, on the radii of the main and shielding coils of the X-, Y-, and Z-gradient coils. Because the X- and Y-gradient coils are bonded onto the Z-gradient coil, it can be assumed that, in an embodiment with grooves, a coaxial groove must also exist, which forms the electrical connection between the two halves of the main gradient coil.
Modern NMR probes are usually manufactured with actively shielded gradient coil systems for generating pulses field gradients. Unlike sensors for magnetic resonance imaging methods (=MRI), most of these gradient coil systems are only uniaxial gradient coil systems, in particular, Z-gradients, in which the most uniform possible gradient of the magnetic field is applied along the z-direction, wherein the z-direction is defined by the direction of the static magnetic field. The effect of this gradient field on a spin I is a rotation about the z-axis through an angle γIGz, wherein G is the gradient amplitude and γI is the gyromagnetic ratio of the spin I. By applying a field gradient, a phase factor of the magnetization encoded along the gradient axis can be induced. In rare cases, gradient coil systems for generating multiple gradient fields are also used, in particular X-, Y-, Z-gradient fields, as are commonly used for MRI.
In nuclear magnetic resonance spectroscopy, it is usual to manufacture probes with integrated pulsed field gradient coils. As a rule, both the probes and the gradient coil systems have a cylindrical and/or hollow cylindrical shape, wherein in particular circularly cylindrical variants are used. These gradient coil systems are usually mounted on (circularly) cylindrical substrates and their conductors substantially occupy the entire lateral surface of the cylinder. In rarer cases and, in particular, for very strong gradient systems that require liquid cooling, the gradient systems are separated from the probes.
There are various manufacturing methods for gradient coil systems: Either they are wound from wire, wherein the wires are usually fixed inside grooves on the formers, or they are cut from, usually metallic, tubes, foils, or conductively coated formers, or manufactured on flexible printed-circuit boards or metal sheets or foils and subsequently mounted on formers.
The gradient turns can be manufactured by two different methods: by the so-called “lane change winding” or “spiral winding” methods. To simplify matters, the following discussion will be limited to Z-gradient coil systems, but applies in the widest sense also to all other gradient coil systems.
In the case of a Z-gradient coil system, in “lane changing” the turns are always located on a z position with the exception of a small section. In the small section, the transition from one z position to the next is executed. In “spiral winding,” the z position is continuously occupied. In particular, wire gradients are usually executed as “lane change windings” because the spiral-shaped grooves cannot be manufactured with high precision or can only be manufactured with high precision with great difficulty. Due to the simpler calculation of the gradient design, however, gradient types manufactured in a different way are usually also manufactured as “lane change windings.”
In order to manufacture a Z-gradient field, a main gradient coil is required, which is usually symmetrical with the xy-plane. To generate the gradient field, however, the direction of rotation of the current must be opposite in the two half spaces. Usually, the two gradient halves are connected in series by means of an electrical connection through the center, wherein this connection is laid on the same radius as the actual gradient turns.
For most NMR applications, actively shielded Z-gradient coil systems are used, wherein due to the necessarily short gradient recovery times, special attention must be paid to shielding the gradients to the outside and their interaction with the magnet and shim system, and with the RF coil systems inside. Actively shielded gradient coil systems usually consist of at least one main gradient coil and one shielding coil each, wherein the shielding coil fully circumferentially surrounds the main gradient coil. In particular, the shielding coils are usually longer than the main gradient coils. Because, for technical reasons, mostly only the lateral surfaces of the cylinder and not the cylinder end faces of the gradient coils are occupied, the missing end faces can be partially compensated for by extending the shielding coils. Moreover, parts of the axial shielding are commonly designed to be on the lateral cylinder surface of the main gradient coil.
Like for MRI, in NMR spectroscopy, the strongest and most efficient field gradients possible are required. In particular, the second point makes it necessary that the radial distance between the main gradient coil and the shielding coil are as large as possible. Because, however, the outer dimensions are determined by the bore of the magnet system, this can only be achieved by reducing the radius of the main gradient coil with respect to the fixed outer radius of the shielding coil.
An NMR probe is not primarily characterized by the gradient coil system that it contains because it is, in particular, designed to transmit and receive RF signals. This is performed with RF coil or resonator systems, which are tuned to the resonance frequencies of the nuclear spins to be measured in a given static magnetic field. There is therefore a lower limit for reducing the radius of the main gradient coils, which is defined by internal volume required for the efficient operation of the RF coil system.
There are basically two possible ways of combining RF coils and gradient coil systems in a NMR probe: Either both systems share the same space, i.e. they are not electromagnetically separated from each other, or the available space is divided into a region for the gradient coil system (gradient region) and a region for the RF system (RF region). In the latter case, an RF shield is placed between the coils and the gradient coil system.
The advantages and disadvantages of the two concepts are as follows: The volume available to an RF system is either not at all or only marginally restricted by a non-shielded gradient coil system as compared with a NMR probe without a gradient coil system. An inserted RF shield substantially reduces the performance of the RF system in some cases because shielding currents have to flow on the RF shield, which, on the one hand, have a dissipative effect and thus impair the Q-factor of the RF system and, on the other hand, generate a field that is opposed to the RF magnetic field and thus reduce the magnetic field amplitude generated per unit current in the measurement volume. Thus, the sensitivity of a NMR probe with RF shielding is reduced as compared with a probe without RF shielding.
However, since a non-RF-shielded gradient coil system in the radio frequency range has a wide spectrum of Eigen-resonances, which, in particular in the case of triple axis gradients, can couple, in some cases massively, with the RF coil system, the use of gradient coil systems without RF shielding is usually very complex or even impossible. Coupling between the Eigen-modes of the gradient and RF coil systems can in some cases cause considerably higher losses in the Q-factor and magnetic field amplitude per unit current than corresponding RF shield would generate.
To avoid this dilemma, the largest possible radii are selected for the gradient coil systems to keep losses resulting from RF shielding as low as possible. However, this results in a lower efficiency of the gradient coil system, which must be compensated for by higher currents and/or higher inductance and higher dissipation during operation.
The object of this invention is therefore to improve an actively shielded gradient coil system of the type described in the introduction by the simplest technical means possible so that the space available in the NMR probe can be divided into an RF region and a gradient region by an RF shield, wherein the volume of the RF region is maximized without loss of performance of the gradient coil system.